High magnetic field electromagnets have become important in various types of equipment over recent years. One important type of such equipment is medical imaging equipment, such as the type commonly referred to as magnetic resonance imaging (MRI) equipment. MRI equipment utilizes the mechanism of nuclear magnetic resonance (NMR) to produce an image, and accordingly imaging systems operating according to this mechanism are also commonly referred to as NMR imaging systems.
As is well known in the field of MRI, a high DC magnetic field is generated to polarize the gyromagnetic atomic nuclei of interest (i.e., those atomic nuclei that have nonzero angular momentum, or nonzero magnetic moment) contained within the volume to be imaged in the subject. The magnitude of this DC magnetic field currently ranges from on the order of 0.15 Tesla to 2.0 Tesla: it is contemplated that larger fields, ranging as high as 4.0 to 6.0 Tesla, may be useful in the future, particularly to perform spectroscopy as well as tomography. The volume of the subject to be imaged (i.e., the volume of interest, or "VOI") is that volume which receives the high DC magnetic field, and within which the DC field is substantially uniform.
Imaging is accomplished in the VOI utilizing the mechanism of nuclear magnetic resonance in the gyromagnetic atomic nuclei contained therewithin. As such, in addition to the large field DC magnet, the MRI apparatus includes an oscillator coil to generate an oscillating magnetic field oriented at an angle relative to the DC field, and at a frequency matching the resonant frequency of the atoms of interest in the selected volume; frequencies of interest in modern MRI are in the radio frequency (RF) range. As the gyromagnetic nuclei in the defined volume will have a common resonant frequency different from atoms outside of the volume, modulation of a gradient magnetic field (produced by a gradient coil) allows sequential imaging of small volumes. The images from the small volumes are then used to form a composite image of the larger volume, such as the internal organ or region of interest. To produce the series of images, the MRI apparatus also includes a detecting coil in which a current can be induced by the nuclear magnetic dipoles in the volume being imaged.
In operation, as is well known, the magnetic dipole moments of those atoms in the volume which are both gyromagnetic and also resonant at the frequency of the oscillating field are rotated from their polarized orientation by the resonant RF oscillation by a known angle, for example 90.degree.. The RF excitation is then removed, and the induced current in the detecting coil is measured over time to determine a decay rate, which corresponds to the quantity of the atoms of interest in the volume being imaged. Incremental sequencing of the imaging process through the selected volume by modulations in the gradient field can provide a series of images of the subject that correspond to the composition of the subject. Conventional MRI has been successful in the imaging of soft tissues, such as internal organs and the like, which are transparent to X-rays.
It is well known in the art that the spatial resolution of MRI tomography improves as the strength of the available magnetic field increases. Conventional MRI equipment useful in diagnostic medical imaging requires high DC magnetic fields, such as 5 kgauss or greater.
Due to the large number of ampere-turns necessary to produce such high magnetic fields, conventional MRI systems now generally utilize superconducting wire in their DC coils. While the magnitude of the current carried in these coils is extremely high, the superconducting material and accompanying cryogenic systems required in such magnets are quite expensive, and also add significantly to the size and weight of the magnet in the MRI apparatus. In the extreme case, some conventional MRI magnets are sufficiently heavy (e.g., on the order of twenty tons) as to limit the installation of the MRI apparatus to a basement or ground floor laboratory. Addition of the necessary coils or iron required to shield the fringe magnetic field generated by such magnets further increases the size, weight and manufacturing costs of the MRI equipment.
By way of background, U.S. Pat. No. 4,783,628 (issued Nov. 8, 1988) and U.S. Pat. No. 4,822,772 (issued Apr. 18, 1989), both incorporated herein by this reference and commonly assigned with this application, describe superferric shielded superconducting magnets. These magnets described in these patents utilize passive shielding of ferromagnetic material, such as iron. The construction of the magnets described in these patents provide a highly efficient magnet, considering the magnetic field strength as a function of the current conducted in the superconducting loops, and with a highly uniform field in the magnet bore even at very strong magnetic fields such as on the order of 4 Tesla; the shielding is also very good in this magnet, with the 5 gauss line at 50 to 100 cm from the outer wall of the bore. The weight and size of the superferric shielded magnets described in U.S. Pat. Nos. 4,783,628 and 4,822,772 can be quite substantial, however, such as on the order of 35 to 130 tons.
Another example of a conventional superconducting magnet, but which relies substantially on active superconducting shielding loops is described in U.S. Pat. No. 4,595,899. The magnet disclosed in this reference has a set of three driving coils surrounded by three shielding coils, with the current through the shielding coils adjusted to exactly cancel the dipole outside of the magnet. External ferromagnetic shielding is also located around the shielding coils to assist in further shielding. Examples of other prior magnets used in MRI are described in U.S. Pat. No. 4,612,505, in which shielding is accomplished by way of magnetic soft iron rods, conducting coils, or both; U.S. Pat. No. 5,012,217, issued Apr. 30, 1992, describes yet another prior superconducting magnet utilizing a combination of active and passive shielding.
While actively shielded magnets greatly reduce the magnet weight relative to superferric shielded magnets, the weight of these magnets is still quite substantial, for example on the order of 20 tons. As a result, when used in medical equipment such as NMR stations, the "footprint" required for installation of the magnet and the weight-bearing capability of the floor of the room are both significant, whether the magnet is constructed of the superferric type, the actively shielded type, or a combination of the two. As a result, from the cost standpoint, it is desirable to reduce the physical size and weight of NMR equipment, to reduce the cost of the NMR laboratory.
In addition to the undesirably large footprint of conventional NMR magnets for medical MRI, a further disadvantage of conventional magnets is patient-related. It has been observed that many patients are uncomfortable when placed in magnets of such length, as the patient's entire body is generally disposed within the magnet during much of the imaging procedure. Indeed, conventional cylindrical NMR magnets have been referred to as "tunnel" magnets, suggestive of the sensation perceived by the human subject when placed inside for an imaging procedure. As an example of the importance of reducing this sensation, U.S. Pat. No. 4,924,185 discloses, relative to another cylindrical superconducting magnet design, that the sense of oppression on the part of the patient is reduced as the ratio of bore length to bore diameter is below 1.90. In addition, especially for patients who are seriously ill, it is essential that the magnet be able to receive the patient without requiring disconnecting life support or monitoring conduits from the patient, and while allowing medical personnel to access the patient during the procedure; conventional cylindrical magnets greatly limit such access.
Accordingly, it is highly desirable to provide NMR tomography equipment of the minimum size but with adequate field strength and uniformity. Toward this end, copending application Ser. No. 715,552, filed Jun. 14, 1991, entitled "A Compact Shielded Superconducting Electromagnet", incorporated herein by this reference and commonly assigned herewith, describes another cylindrical superconducting magnet which advantageously uses a combination of superferric shielding outside of the shielding coils. The magnet disclosed therein thus can be shorter in length while still providing high DC field in the bore and low fringe field away from the magnet.
Copending application Ser. No. 869,544, filed Apr. 15, 1992 in the name of Sergio Pissanetzky, and entitled "An Ultrashort Cylindrical Shielded Electromagnet For Magnetic Resonance Imaging", commonly assigned herewith, incorporated hereinabove by reference, and being the parent of this continuation-in-part application, discloses a short cylindrical superconducting magnet useful in NMR imaging equipment. This application further discloses an important methodology used to design the coils in the DC field generating magnet in such a manner as to optimize the strength and uniformity of the field in the volume of interest (VOI) while maintaining low fringe fields away from the bore. As will be described in further detail hereinbelow, this methodology also proves useful in the design of magnets of other shapes.
By way of further background, U.S. Pat. No. 4,689,591 discloses a superconducting magnet having a plurality of coaxial coils arranged asymmetrically along the axis, resulting in a volume of interest that is offset from the midplane of the magnet. The volume of interest in this magnet, while offset, remains deeply within a cylindrical bore, however, requiring whole body insertion of the patient for MRI procedures.
Another known type of conventional MRI magnet is of the Helmholtz coil type, including such a magnet which utilizes thermally insulated niobium/tin superconducting material. However, this magnet also requires the patient's whole body to be inserted between the Helmholtz coils.
By way of further background, U.S. Pat. No. 5,049,848, issued Sep. 17, 1991, discloses a magnet configuration suited for MRI in mammography. This magnet configuration is of rectangular shape, and includes permanent magnets (5, 6, 7, 8) for generating magnetic flux in two planes in the gap g within which the imaging is to take place. A shimming electromagnet (14) is disclosed as being placed behind the patient, for reducing front edge fringe field.
By way of further background, notched cylindrical coil systems for providing strong magnetic fields are known, as described in M. W. Garrett, "Thick Cylindrical Coil Systems for Strong Magnetic Fields with Field or Gradient Homogeneities of the 6th to 20th Order", J. Appl. Phys., Vol. 38, No 6 (1967), pp. 2563-86. As described in this reference at pages 2578-2583, expansion of ideal magnet elements to larger cross sections by modifying the geometry in an iterative fashion according to Lyle's Principle can be used to arrive at a magnet having a negative current polarity notch or cavity coil within the otherwise cylindrical positive coil; the notch may be at either the inner or the outer radial surface (see FIG. 2 of the reference), or even wholly within the positive coil.
In the method described in the above-cited Garrett article and also according to other conventional methods for designing cylindrical magnets, the designer relies on the property that the axial component of the magnetic field in the bore is a harmonic function within the volume of interest (VOI), which can be expanded into a series of spherical harmonics. The coefficients of this expansion may be expressed as axial derivatives of the axial magnetic field at the origin (center of the VOI). If one assumes that the current density in the cross-section of each magnet coil is constant, these axial derivatives may be calculated directly from coil geometry, without requiring integration of the Biot-Savart Law, allowing the geometry of the magnet to be adjusted so that the undesirable harmonics of the axial field in the VOI vanish. According to this generalized technique, computer-ready methods have been developed as have tabular design conditions (see the Garrett article cited above), facilitating the design of such magnets. This general method constrains the location of the VOI to a high degree, however, in order for the calculations to be readily performed; as such, this general method is practically applicable for a VOI centered at the midplane of the cylindrical magnet.
It is an object of the present invention to provide a method of fabricating an electromagnet for optimizing the field strength and uniformity in a selected volume of interest.
It is a further object of the present invention to provide such a method of fabricating a magnet which is applicable to magnets of various symmetry, including cylindrical and planar magnets.
It is a further object of the present invention to provide a magnet constructed according to such a method.
It is a further object of the present invention to provide such a magnet which allows for optimization of the design for a volume of interest which is not necessarily centered within the magnet bore.
It is a further object of the present invention to provide such a method which allows the center of the volume of interest to be located outside of the magnet bore.
It is a further object of the present invention to provide a superconducting magnet for use in NMR equipment which does not require insertion of the body of the patient into the magnet bore.
It is a further object of the present invention to provide such a magnet which is of sufficient field strength to enable in vivo NMR tomography of the internal organs of humans.
It is a further object of the present invention to provide such a magnet which is suitable for the image of specific organs, such as the brain, the female breast, and the like.
It is a further object of the present invention to provide such a magnet which makes the NMR tomography equipment substantially portable.
Other objects and advantages of the present invention will be apparent to those of ordinary skill in the art having reference to the following specification together with the drawings.